Field of the Invention
The present invention concerns a method to generate magnetic resonance image data of an examination subject with the use of a magnetic resonance imaging system, and a magnetic resonance imaging system for implementing such a method.
Description of the Prior Art
Imaging systems that are based on magnetic resonance measurement (of nuclear spins), known as magnetic resonance tomography systems have been successfully established and proven through numerous applications. In this type of image acquisition, a static basic magnetic field B0 that serves for initial alignment and homogenization of magnetic dipoles to be examined is superimposed with a rapidly switched magnetic field (known as the gradient magnetic field) for spatial resolution of the imaging signal. To determine the material properties of an examination subject to be imaged, the dephasing or relaxation time after a deflection of the magnetization out of the initial alignment is determined so that different relaxation mechanisms or relaxation times typical to the material can be identified. The deflection typically takes place by radiating a number of RF pulses, and the spatial resolution is based on a temporally fixed manipulation of the deflected magnetization with the use of the gradient magnetic field by activation of gradient pulses in a measurement sequence, which establishes a precise chronological sequence of RF pulses, gradient pulses and the acquisition of measurement values.
If a switching sequence of the gradient magnetic field in a measurement sequence is delayed relative to an expected point in time of the switching, this leads to inaccuracies in the spatial resolution of the magnetic resonance signal that cause distortions and other errors in the magnetic resonance imaging of an examination subject. This delay is designated as a “switching lag” in the following.
An association between measured magnetization (from which the mentioned material properties can be derived) and a spatial coordinate of the measured magnetization typically takes place with the aid of an intermediate step. In this intermediate step, acquired raw magnetic resonance data are entered into a memory in an organization known as “k-space”, wherein the coordinates of k-space are coded as a function of the gradient magnetic field. The gradient magnetic field varies the resonance frequency (Larmor frequency) and, for example, also the phase position of the magnetization deflected by an RF pulse in a spatially dependent manner, such that spatial information in k-space is obtained by the designation of phase position and resonance frequency of the measured magnetization. Thus k-space is also known as the frequency domain. In other words, spatial information is based, with phase coding and frequency coding, on the coordinate system of k-space (spatial frequency) and is defined as a function of the gradient magnetic field. The magnitude of the magnetization (in particular of the transverse magnetization, defined in a plane transversal to the previously described basic magnetic field) at a defined location of the examination subject can be determined from the readout point in k-space with the use of a Fourier transformation, with which the signal strength of the signal in the spatial domain can be calculated from the signal strength (magnitude of the magnetization) that is associated with a defined frequency (the spatial frequency).
K-space thus forms an inverse Fourier space relative to positional space of the examination subject, such that the magnetic resonance signals are transformed into the positional space with the aid of a Fourier transformation to create the magnetic resonance image. The gradient magnetic field thus determines a point in k-space, wherein the time curve of the gradient magnetic field establishes a series of k-space points, which can be designated as what is known as the “trajectory” through k-space or also as a “projection”.
The aforementioned switching lag can disadvantageously reach an order of microseconds in present magnetic resonance imaging systems, and therefore markedly exceed the switching delay of an RF pulse for the deflection of the magnetization. If this is the case, the gradient magnetic field assumes a value other than the one expected at a readout point of the raw magnetic resonance data, and a gradient magnetic field or a phase position of the spins that corresponds to an expected k-space coordinate is only achieved at a later point in time. It results from this that the measured magnetic resonance signal is associated with a displaced coordinate in k-space, since the gradient magnetic field or the required phase position of the spins does not have the expected value at the point in time of measurement.
If a displacement of the k-space coordinates of the trajectory occurs due to the switching lag so that an approximately coherent displacement (explained in detail later) is present for all trajectories—for example given line-by-line Cartesian scanning of k-space—the switching lag has nearly no effect on the quality of the imaging of the examination subject since the additionally created phase or frequency shift is the same for all k-space points. However, if this is not the case—for example if the scanning of k-space is selected in a particular path through k-space and takes place radially, for example—this inevitably leads to severe artifacts in the imaging. In this case, a correction of the k-space points of the trajectories should take place in order to be able to implement a transformation of the magnetic resonance signals into positional space of the examination subject while avoiding distortions and image artifacts.
The switching lag is particularly problematic given image data that are determined with the use of known “UTE sequences” (U=ultrashort, TE=echo time). An echo signal of the magnetization which should be detected thereby occurs after an “ultrashort time” (i.e. after an echo time TE1 of between 5 μs and 200 μs) after emission of an RF pulse. The detection of magnetic resonance signals given the UTE sequences therefore typically takes place with trajectories that travel outward radially from the origin point of k-space to the boundaries of scanned k-space (what are known as “half projections”), since the echo time TE1 is not sufficient in order to start a trajectory from a border region of k-space, for example. The switching delay given the acquisition of half projections is disadvantageously such that significant portions of the magnetic resonance signals can only be imprecisely associated with a spatial frequency due to the short time sequence between excitation and signal acquisition. The switching lag thus leads to particularly severe artifacts in image reconstruction.
In order to improve the association of magnetic resonance signals with spatial frequencies, a method for correction is known that corrects the k-space points of the trajectories with the aid of a correction value. The correction value is added to the correction method via input, wherein the input can take place manually or also from a database, for example. However, this method is time-consuming or based on general models for the displacement due to the switching lag, such that significant effort arises in the acquisition of the magnetic resonance imaging, and moreover the correction no longer takes place optimally.